1. Field of the Invention
The present invention concerns a method for spatially resolved determination of a magnetic resonance (MR) parameter, in particular an MR parameter that influences an MR signal detected given an MR measurement of a region of an examination subject. Furthermore, the invention concerns a magnetic resonance system for implementation of such a method.
2. Description of the Prior Art
Magnetic resonance tomography (MRT) is an imaging modality that enables the acquisition of two-dimensional or three-dimensional image data sets that can depict structures inside an examination subject (in particular even soft tissue) at high resolution. In MRT, protons in the examination subject are aligned in a basic magnetic field (B0) so that a macroscopic magnetization appears that is subsequently excited via the radiation of RF (radio-frequency) pulses. The decay of the excited magnetization is subsequently detected by means of one or more induction coils, wherein a spatial coding of the acquired signal is achieved via the switching of slice selection, phase coding and frequency coding gradients before or, respectively, during the acquisition. The acquisition of the decay signals thereby regularly takes place with a quadrature detection so that both the phase and the amplitude of the signal are detected. The signals detected in positional frequency space (frequency domain) (k-space) can accordingly be represented as complex numbers and be transformed by means of a Fourier transformation into image space (image domain), in which phases and magnitudes can now be determined with spatial resolution.
In many imaging methods only the magnitude of the complex image data is used to create an intensity image. The phase information is discarded. Furthermore, a combination of the magnitude data that were acquired with different coils is not optimal with regard to the signal-to-noise ratio (SNR).
For example, in conventional T2* (observed spin-spin relaxation time) or R2* (01/T2*) imaging the magnitudes of three or more images are considered that were acquired at different echo times (TEs) from an individual proton species (for example via the use of a fat suppression). The T2* time can subsequently be determined with spatial resolution via the adaptation of a decay function to the magnitudes in the image data. However, this approach is very time-consuming and correspondingly prone to movement artifacts.
Other imaging methods use the acquired phase information. For example, the differences of the magnetic susceptibility of different tissue lead to phase differences. In susceptibility-weighted imaging (SWI), an expanded contrast signal image is generated from the acquired magnitude and phase information. This expanded contrast signal image has a contrast that is dependent on the oxygen content of the blood. Additional examples generally include phase contrast imaging as well as proton resonance frequency (PRF) shift thermometry. In this type of imaging, a shift of the phase in acquired phase images is detected that is caused by a shift of the proton resonance frequency due to a temperature change. In general a shift of the phase is based on a shift of the resonance frequency of the excited protons.
In addition to these phase shifts with information content, there is a series of effects that cause unwanted phase shifts and can conceal usable information. Among these effects are an inhomogeneity of the static B0 field, the susceptibility of articles and materials within or in proximity to the patient, phase shifts of the radiated RF pulses and errors in the chronology of the acquisition sequence. Phase shifts that can develop differently for different acquisition or reception coils also can occur in the acquisition chain or in the acquisition channel of the respective coil.
These phase shifts make it difficult to compare and combine image data acquired at different echo times with one another. In particular, the combination of MR data acquired with different acquisition or reception coils while acquiring phase information has proven to be difficult, since each acquisition channel has a different phase shift. Objects within the examination subject—for example air bubbles, implants, needles or the like—can also lead to susceptibility artifacts, and thus also to phase shifts.
It is desirable to combine MR parameters, for example the T2* relaxation time or the resonance frequency shift ω, as precisely and effectively as possible. Magnetic resonance (MR) data acquired for different echo times or with different reception coils should be optimally combined with one another in the determination so that the signal-to-noise ratio is improved and usable phase information is retained. In order to enable shorter scan durations, the method should also be capable of enabling such a combination for accelerated acquisition methods and multi-echo imaging sequences. Moreover, the data should be combined in a well-defined manner in order to be able to make reasonable statements about the data acquired in such a manner, for example about errors or noises.
A method known from the prior art for R2* imaging is k-TE GRAPPA, which is described in detail in “k-TE Generalized Autocalibrating Partially Parallel Acquisition [sic] (GRAPPA) für Accelerated Multiple Gradient-Recalled Echo (MGRE)R2* Mapping in the Abdomen”, by Xiaoming Yin et al., Magnetic Resonance in Medicine 61:507-516 (2009). The method uses a partially parallel imaging method (GRAPPA) in combination with a view-sharing method in which missing k-space lines in incompletely scanned k-space are reconstructed on the basis of k-space lines acquired with adjacent coils and temporarily adjacent sequences. The result of the method is a series of images of different echo times (TE), wherein the image noise in the image data varies spatially due to the reconstruction process. How the acquired image data can be combined optimally with regard to SNR is not disclosed in this publication.
Given a combination of image data that takes phase into account, in conventional methods the phases of adjacent image points are compared in order to produce a total phase estimation. Image points with large phase variation, for example in regions with low SNR or along tissue boundaries, can interfere with the phase correction method. Furthermore, methods are known from U.S. Pat. No. 7,227,359 B2 (for example) that are based on phase gradients in the image data and that implement a region expansion (region growing) to determine the phases using a seed image point.
For image data that were acquired with a multi-echo imaging sequence, for example with a single-shot or segmented EPI (echoplanar imaging) sequence, a “characteristic” echo time (TE) is normally associated with the data. This characteristic TE is typically the TE with which central k-space lines were scanned. However, in such sequences different spatial frequencies (k-space lines) are scanned using different TEs, such that these respectively contribute to an error in the reconstructed image data depending on the respective TE. The errors in the phase accordingly depend on the acquisition sequence and moreover have a spatial dependency, such that they can only be predicted with difficulty.